C08F234/00—Copolymers of cyclic compounds having no unsaturated aliphatic radicals in a side chain and having one or more carbon-to-carbon double bonds in aheterocyclic ring

C08F234/02—Copolymers of cyclic compounds having no unsaturated aliphatic radicals in a side chain and having one or more carbon-to-carbon double bonds in aheterocyclic ring in a ring containing oxygen

G21K5/10—Irradiation devices with provision for relative movement of beam source and object to be irradiated

Description

BackgroundField of the Invention

This invention relates generally to radiation therapy
equipment for the treatment of tumors, or the like, and
specifically to a mechanism for regulating the dose of
radiation within irregularly shaped zones within a patient
and to a mechanism for verifying radiation intensity
directed at, and the dose of radiation absorbed within,
irregularly shaped zones of a patient.

Description of the Art

Medical equipment for radiation therapy treats
tumorous tissue with high energy radiation. The dose and
the placement of the dose must be accurately controlled to
insure both that the tumor receives sufficient radiation to
be destroyed, and that damage to surrounding and adjacent
non-tumorous tissue is minimized.

Internal-source radiation therapy places capsules of
radioactive material inside the patient in proximity to the
tumorous tissue. Dose and placement are accurately
controlled by the physical positioning of the isotope.
However, internal-source radiation therapy has the
disadvantages of any surgically invasive procedure,
including discomfort to the patient and risk of infection.

External-source radiation therapy uses a radiation
source that is external to the patient, typically either a
radioisotope, such as 60Co, or a high energy x-ray source,
such as a linear accelerator. The external source produces
a collimated beam directed into the patient to the tumor
site. External-source radiation therapy avoids some of the
problems of internal-source radiation therapy, but it
undesirably and necessarily irradiates a significant volume
of non-tumorous or healthy tissue in the path of the
radiation beam along with the tumorous tissue.

The adverse effect of irradiation of healthy tissue
may be reduced, while maintaining a given dose of radiation
in the tumorous tissue, by projecting the external
radiation beam into the patient at a variety of "gantry"
angles with the beams converging on the tumor site. The
particular volume elements of healthy tissue, along the
path of the radiation beam, change, reducing the total dose
to each such element of healthy tissue during the entire
treatment.

The irradiation of healthy tissue also may be reduced
by tightly collimating the radiation beam to the general
cross section of the tumor taken perpendicular to the axis
of the radiation beam. Numerous systems exist for
producing such a circumferential collimation, some of which
use multiple sliding shutters which, piecewise, may
generate a radio-opaque mask of arbitrary outline.

As part of collimating the beam to the outline of the
tumor, the offset of the radiation beam, with respect to a
radius line between the radiation source and the center of
rotation of the radiation source, may be adjusted to allow
the treated area to be other than at the center of
rotation. Simultaneously changing the offset and the width
of the radiation beam as a function of gantry angle allows
tumorous tissue having an irregular cross-section within a
plane parallel to the radiation beam to be accurately
targeted. The width and offset of the radiation beam may
be controlled by the use of a multiple-leaf circumferential
collimator.

Adjustment of the offset and size of the radiation
beam at various gantry angles allows considerable latitude
in controlling the dose. Nevertheless, even using these
techniques, there is still a considerable amount of
undesired dose imparted to healthy tissue, especially where
a treatment volume is concave or highly irregular within
the plane parallel to the radiation beam.

A radiotherapy machine providing much reduced
irradiation of healthy tissue is presented in U.S. Patent
number 5,317,616 issued to Stuart Swerdloff et al, on May
31, 1994. The architecture
described in the above application employs a radiation
source constrained to rotate within a single gantry plane
about a patient while intensity of individual rays of the
radiation beam are modulated by a set of opaque leaves
which move into and out of the radiation beam.

The leaves slide into the radiation beam in a closed
state and out of the radiation beam in an open state to
allow unobstructed passage of a given ray of the beam. By
employing appropriate planning techniques the dose absorbed
by each slice of the tumor may be controlled to irradiate
even tumors having a concave cross section within the
gantry plane. This ability to control not just the outline
of the radiation but the intensity of each individual ray
allows extremely precise control of the irradiation volume.

An entire tumorous volume may be treated by moving the
patient with respect to the gantry plane and irradiating
each tumor slice separately. The compensator and gantry
configuration together substantially increase the ability
to conform a radiation dose to an arbitrarily shaped tumor
while employing a simplified two-dimensional protocol.

Besides simplifying the irradiation protocol, a single
plane configuration provides a number of other benefits.
These benefits include the ability to use a single ring
gantry to support both a radiation source and a CT system,
the advantage of reduced interference between the radiation
source and patient (or table) and the advantage of
simplified shielding requirements. Furthermore, the
simplified architecture enables a therapist to employ a
helical scanning method to smoothly irradiate along the
length of a tumor so as to avoid irradiation hot spots or
irradiation gaps.

Despite the advantages of constraining a therapy
machine to operation with a single plane, a single plane
machine presents various problems.

First, there is the need for dose verification. The
destructive potential of a radiation beam to healthy tissue
and the necessity of insuring that tumorous tissue receives
sufficient radiation makes treatment verification a
required part of radiation therapy. With conventional
therapy machines, films may be exposed during a therapy
session both to confirm the location of an irradiated area
and to provide a record of radiation dose. Because the
radiation source employed by the above described gantry
configuration constantly sweeps around a gantry to produce
a moving beam, a film verification system employing
traditional film movement techniques would be unworkable.
The verification problem is more profound if scanning is
performed in a helical manner.

In addition, a single plane system must employ a
compensator capable of varying the intensity of individual
rays of a beam in order to properly treat a tumor. The
reliability of such a compensator must be extremely high --even
a single improperly attenuated beam ray passing
through the compensator undetected and irradiating healthy
tissue due to a failed compensator component could result
in severe damage to healthy tissue.

In practice, regardless of accurate machining
techniques, small gaps between the leaves in the
compensator described above, that allow unattenuated rays
to pass, are needed to prevent frictional contact. The
exacting tolerances necessary to minimize the size of the
leaf gaps are costly and such high tolerance components are
prone to failure.

Another problem with the compensator described above
is that the rays are not uniformly attenuated by the moving
leaves because the leaves cannot be moved instantaneously.
A leaf initially occludes the entire depth of its
associated ray within the beam. As the leaf begins to move
out of the beam, part of the ray is occluded and another
part is left unobstructed. Eventually, the entire ray is
unobstructed. The same non-uniform beam attenuation is
again encountered when the leaf moves back into the beam.

This gradation in attenuation can be minimized by
equipping the compensator with more powerful actuators to
drive the leaves in and out of the fan beam width more
rapidly. Bigger actuators, however, are more costly to
utilize and maintain. Alternatively thinner leaves may
be used which are light weight so that they can be moved
more quickly. This, however, demands more actuators,
creates more leaf gaps, and must be accommodated by a
much more complicated control system.

EP-A-0259989 discloses a radiation therapy machine
in which a radiation beam is generated which is to be
directed towards the patient. In order to produce
irregular field shapes, an attenuation means is provided
in the form of a multileaf collimator. The collimator
has a multiplicity of heavy metal bar leaves which divide
the beam at adjacent rays. A compensator is attached to
each leaf, enabling the local intensity of each ray to
be controlled. The compensators may be controlled to
control the attenuation. Thus their document
corresponds to the pre-characterising part of claim 1.
However, EP-A-0259989 suggests that the attenuation
control should be achieved by monitoring the rays which
have pierced the patient.

EP-A-0371303 also discloses a radiation therapy
machine in which a multileaf collimator divides the beam
into adjacent rays, and the local intensity of the rays
controlled. Again, the attenuation control is not
discussed in detail.

The present invention seeks to provide a
verification system that can be used in conjunction with
a radiation intensity compensator to minimize the
possibility of an uncontrolled beam ray irradiating non-tumorous
tissue. In one embodiment, the verification
system may collect tomographic data on absorbed radiation
within the patient and generate tomographic absorption
images therefrom. These images may be used for radiation
dose verification as well as for planning subsequent
therapy sessions.

According to the present invention there is
disclosed a radiation therapy machine having a radiation
source for producing a radiation beam directed toward a
patient at a gantry angle, the beam including a plurality
of adjacent rays, and a compensator having an attenuation
means disposed between the radiation source and the
patient for independently controlling the fluence of each
ray of the beam, and a compensator control for
controlling the attenuation means according to first
desired ray fluences at a first gantry angle;
wherein:
the compensator further includes a pre-patient
monitor disposed between the attenuating means and the
patient having a plurality of monitor segments, adjacent
monitor segments subtended by adjacent rays of the beam,
each segment producing a fluence signal proportional to
the measured fluence of the ray subtending the segment,
comparison means for comparing the desired fluence of a
ray to the measured fluence to produce a difference
value, and limit means for producing an error signal if
the difference value is outside of a predetermined error
range.
controlled by the compensator. Large discrepancies between
the measured and desired signals indicate a malfunction.

A high signal may also be produced indicating that the
measured signal is too high relative to the desired signal
yet within the predetermined error range and a low signal
may be produced indicating that the measured signal is low
relative to the desired signal yet within the predetermined
error range. The compensator control receiving these
signals may adjust the attenuation means to reduce a second
ray fluence at a second gantry angle in response to
receiving a high signal and to increase a secondary ray
fluence at a second gantry angle in response to a low
signal.

The system
may compensate for deviations in ray fluence by
adjusting ray fluence at later angles. The effect of the
difference between a desired signal and the measured signal
can be reduced as the radiation from the two angles
combine.

The attenuating means may be a
plurality of radiation attenuating leaves in a support
structure positioned generally between the radiation source
and the patient. The support structure guides the leaves
between a closed state within the radiation beam, each leaf
thus occluding one ray of the beam, and an open state
outside of the radiation beam to allow unobstructed passage
of the ray.

A motivator independently moves each leaf between the
open and closed states to effect an open-to-closed desired
ratio producing the fluence of each ray. The motivator may
be a first set of actuators connected by linkages to
individual leaves to move the leaves with the movement of
the armatures.

A position sensor determines when each leaf is in the
open state and when each leaf is in the closed state and
produces an actual ratio of the period of time the leaf is
in the open state to the period of time the leaf is in the
closed state. An error detector generates an error signal
by comparing the actual ratio to the desired ratio.

Another object of the invention is to provide a
mechanical compensator system with extremely high
reliability. The error signal generated by the error
detector may be used to indicate a malfunction.

In another embodiment, the supporting structure may
guide the leaves between a closed state centered within the
radiation beam, the leaf thus occluding one ray of the
beam, a first open state with the leaf displaced outside of
the radiation beam on a first side of the one ray, and a
second open state with the leaf displaced outside of the
radiation beam on a second side of the one ray. In this
embodiment the motivation means moves each leaf alternately
between the first open and closed state and between the
closed and second open states.

The compensator can attenuate individual rays of a high
energy fan beam to provide uniform attenuation across the
thickness of the fan beam. Provided the acceleration and
velocity of the leaf in each direction is identical, the
uneven radiation caused by motion of the leaf in one
direction is canceled by the radiation caused by the motion
of the leaf in the second direction.

In yet another embodiment, the plurality of
attenuating leaves in the attenuating means above includes
a first plurality of radiation attenuating leaves and a
second plurality of radiation attenuating leaves and the
supporting structure includes a first supporting structure
for guiding the first plurality of leaves between a closed
state, each leaf thus occluding every other ray of the
beam, and an open state outside of the radiation beam and a
second support structure for guiding the second plurality
of leaves between a closed state, each leaf thus occluding
those rays of the beam not occluded by the first plurality
of leaves when the latter are in the closed state, and an
open state outside of the radiation beam wherein the first
plurality of leaves in the closed state are positioned
closer to the radiation source than the second plurality of
leaves in the closed state.

A post-patient monitor may be disposed opposite the
pre-patient monitor with respect to the patient and within
the fan beam for determining a post-patient fluence of each
ray of the beam exiting the patient. An absorption
calculator compares the pre-patient fluence to the post-patient
fluence to produce an absorption value for each
ray, the absorption values together providing an absorption
profile for the fan beam at a given gantry angle.

The foregoing and other objects and advantages of the
invention will appear from the following description. In
the description, reference is made to the accompanying
drawings which form a part hereof and in which there is
shown by way of illustration several preferred embodiments
of the invention. Such embodiments do not necessarily
represent the full scope of the invention, however, and
reference must be made therefore to the claims herein for
interpreting the scope of the invention.

Brief Description Of The Drawings

Fig. 1 is a perspective view of the compensator
assembly used in the present invention, showing the
compensator leaves and their associated electromagnetic
actuators;

Fig. 2 is a cross section of the compensator assembly
of Fig. 1 along line 2-2 showing the trapezoidal aspect of
each compensator leaf, for a fan beam of radiation, and the
guide rails for supporting the compensator leaves when they
move;

Fig. 3 is a cutaway perspective view of a set of guide
rails and one leaf of Fig. 2 showing a collar for
supporting the leaf;

Fig. 4 is a plan view of a portion of the mounting
rack and one leaf of Fig. 1 showing a trigger bore, a
trigger and a light emitting diode and light detector pair
with the leaf in its fully closed position;

Fig. 5 is a block diagram showing the elements of a
radiation therapy apparatus incorporating a conventional CT
scanner and the compensator of the present invention and
including a computer suitable for controlling that
compensator per the present invention;

Fig. 7 is a diagrammatic representation of a patient
receiving radiation therapy, showing the scatter kernel and
the coordinate system used to describe the present
invention;

Fig. 8 is a perspective representation of a monodirectional
scatter kernel associated with a radiation beam
at one gantry angle;

Fig. 9 is a perspective representation of a composite
multidirectional scatter kernel associated with a plurality
of radiation beams at multiple gantry angles;

Fig. 10 is a block diagram depicting the fluence
profile calculator which takes a desired dose map and
calculates a fluence profile;

Fig. 11 is a block diagram depicting the overall
iterative method of controlling the compensator of the
present invention, employing the fluence profile
calculation method of Fig. 10;

Figs. 12(a)-(c) are perspective views of plots showing
the error between the desired dose distribution and the
actual dose distribution obtained with the present
invention for one, two and four steps of iteration
respectively.

Figs. 13(a)-(c) are schematic views showing the
relationship between an irradiation window and adjacent
tumor slices as the radiation source rotates about the
gantry from 0° to 90° to 180°.

Fig. 14 is a schematic view showing the general
orientation of the monitoring chambers in relation to the
compensator and the patient.

Fig. 16 is a perspective view of a second embodiment
of assembly used in the present invention showing two
levels of the compensator leaves, their associate pneumatic
cylinders and their associated solenoid assemblies;

Fig. 17 is a cross-section of the compensator assembly
of Fig. 16 along line 17-17 showing both levels of
compensator leaves, the trapezoidal aspect of each
compensator leaf, the associated solenoid stopping
assemblies, and the guide rails for supporting the
compensator leaves when they move;

Fig. 18 is a cutaway perspective view of a set of
guide rails and one leaf from the first level and one leaf
from the second level of Fig. 16;

Figs. 19(a)-(c) are side views of a leaf showing the
relationship between a leaf, the beam and the solenoid
assemblies when the leaf is in the first open Fig. 19(a),
closed Fig. 19(b) and second open Fig. 19(c) states;

Fig. 20 is a detailed cross sectional view of the
solenoid stop assembly shown in Fig. 17 taken along the
line 20-20;

Fig. 21(a)-(g) are graphs showing the changing fluence
gradient across the beam as a leaf is moved from a closed
state in Fig. 21(a) into a first open state in Fig. 21(c),
back into the closed state in Fig. 21(e), into a second
open state and again back into the closed state Fig. 21(g).

Detailed Description Of The Preferred Embodiment

Referring to Fig. 1, a radiation therapy unit 10
suitable for use with the present invention includes a
radiation source 12 producing a generally conical radiation
beam 14' emanating from a focal spot 18 and directed
towards a patient 17 (not shown in Fig. 1). The conical
beam 14' is collimated by a radiation opaque mask 16
constructed of a set of rectangular collimator blades to
form a generally planar fan beam 14 centered about a fan
beam plane 20.

I. The Compensator

A compensator 22 is centered in the fan beam 14 and
about the fan beam plane 20, prior to the radiation being
received by the patient 17, and includes a plurality of
adjacent trapezoidal leaves 30 which together form an arc
of constant radius about the focal spot 18. The leaves 30
are held in sleeves 24. The sleeves 24 are constructed of
radio-translucent materials and attached at their inner
ends 23 to a mounting plate 26 which is fixed relative to
the focal spot 18. The mounting plate 26 is constructed of
a sturdy, radiopaque material and is positioned just
outside the fan beam 14 to prevent interference with the
fan beam 14.

Preferably, the leaves 30 of the compensator 22
subtend the entire fan beam 14 to divide the fan beam 14
into a set of adjacent slab-like rays 28 at offset angles
1. Referring also to Fig. 2, each sleeve 24 is open at its
outer end 27 to receive, by sliding, a comparably sized
trapezoidal leaf 30 constructed of a dense, radiopaque
material such as lead, tungsten, cerium, tantalum or a
related alloy.

Each leaf 30 may slide completely within its
corresponding sleeve 24 to block the ray 28 associated with
that sleeve 24. When the leaf 30 blocks its corresponding
ray 28, it is referred to as being in a "closed state".
The sleeves 24 are of ample length to permit each leaf 30
to slide out of the path of the fan beam 14, so as to leave
its corresponding ray 28 completely unobstructed, and yet
to still be guided by the sleeve 24. In this non-blocking
position, a leaf is referred to as being in the "open
state".

Each leaf 30 moves rapidly between its open and closed
states by means of a primary corresponding relay-like
electromagnetic actuator 32 connected to the leaf 30 by a
slider member 34. The actuators 32 have internal armatures
(not shown) received within solenoid electromagnets. The
armature may be moved at high velocity by means of varying
electrical excitations of their associated electromagnets.
The electrical excitations are provided by a compensator
control (not shown in Figs. 1 or 2) to be described below.
The actuators 32 are capable of applying high forces to the
leaves 30 to move them rapidly and independently between
the open and closed states.

Each leaf 30 is also provided with a back up actuator
35 located below its primary actuator 32 on the outer edge
of the leaf 30. The secondary actuator 35 is employed when
the primary actuator 32 fails as will be described in more
detail below.

Referring now to Figs. 2 and 3, the leaves 30 are
supported and guided within the sleeves 24 by guide rails
36 fitted into notches 38 cut along the edges of the leaves
30. The notches 38 allow the guide rails 36 to slidably
retain the leaves 30 within the sleeves 24 during motion
between the open and closed states.

In the closed state, the inner end 40 of each leaf 30
is captured by a rigid collar 42 attached to the mounting
plate, which aligns the leaf 30, more accurately than may
be done by the guide rails 36, with the mounting plate 26
and hence with the fan beam 14. Whereas the guide rails
36, which are ideally radio-translucent, are relatively
insubstantial, in contrast, the collar 42, positioned
outside the fan beam 14 on the mounting plate 26, need not
be radio-translucent and hence is more substantial in
construction. A collar (not shown) similar to collar 42,
supports each leaf 30 when it is fully in the open state.
Because the leaves 30 spend most of their time fully in the
open or closed states, they are, at most times, firmly
located by a supporting collar 42.

Of concern is the reliability of leaf 30 switching
because one inoperable leaf 30 would prevent treatment.
The number of switching cycles S (opening and closing) in a
year is given by the following equation:
S=250ρLBΔz
where 250 is typical of the number of treatment days per
year in North America and P is the number of patients
treated per day. S is going to be on the order of millions
of cycles per year.

If the probability of a single leaf 30 failing in a
year is P1 then the probability of any leaf within the
system failing P is given by:
P=1-(1-P1)n
where n is the number of leaves 30. Since the leaves 30
are light, the slide member 34 is under little stress and
so it has a low probability Pslide of failure.

Referring now to Figs. 3 and 4, a light opaque trigger
member 19 is integrally attached to the inner edge 40 of
each leaf 30, the trigger member 19 extending laterally
outward parallel to the movement axis 25 of each leaf 30.
When the leaf 30 is in the closed state (see Fig. 4), the
trigger member 19 passes through a trigger bore 31 in the
back wall 29 of the mounting plate 26.

A plurality of light emitting diode 21 and light
detector 15 pairs are positioned on the outer surface of
the back wall 29, the elements of each pair opposing each
other on opposite sides of an associated trigger bore 31.
When a leaf 30 is in the closed state, its associated
trigger member 19 extends through the trigger bore 31 and
blocks the light path 33 between the light emitting diode
21 and light detector 15. Referring to Fig. 3, when the
leaf 30 is not in the closed state, the trigger member 19
does not extend through the trigger bore 31 and hence light
passes from the emitting diode 21 to the light detector 15.
The probability of the light detector 15 failing is Pverify.

An axially compressible spring 39 is provided between
the outer edge 41 of each leaf 30 and the front wall (not
shown) of the mounting plate 26 to move the leaf 30 to the
closed position absent force from its actuator 32. A
system employing a light detector 15, primary and backup
actuators 32, 35 and which are regularly serviced, would
have downtime due to failure of a leaf 30 according to the
probability equation:
P1=1-(1-Pslide)(1-Pverify)[1-(PactuatorNmaint)2]
wherein Pactuator is the probability of the actuator or
control electronics failing during a year and Nmaint the
number of maintenance services in the year (e.g. it could
be every day). The exponent 2 accounts for the failure of
both actuators in the maintenance period. This equation
basically says that failure of a leaf 30 will happen if the
slider 34 fails, the verification system fails or both of
the actuators 32, 35 fail before the system can be
serviced.

Referring to Figs. 5 and 14, a pre-patient multi-segment
ion chamber 47 is positioned between the
compensator 22 and the patient 17. Each ray 28 of the fan
beam 14 subtends a separate monitor segment 49 as it passes
through the ion chamber 47. A post-patient multi-segment
ion chamber 53 is positioned directly opposite the
radiation source 12 on the gantry 44 so as to intercept the
fan beam 14 as it exits the patient 17. The separate
monitor segments 54 of the second ion chamber 53, like the
monitor segments 49 of the first ion chamber 47, are each
subtended by individual rays 28 of the fan beam 14. The
ion chambers 47, 53 produce signals indicating the fluence
of rays 28 (as generally understood in the art) and as used
by the computer 51 to determine radiation dose in a manner
to be described below.

II. Radiation Therapy Hardware

Referring now to Fig. 5, the radiation source 12 is
mounted on a gantry 44, the latter rotating within the fan
beam plane 20 about a center of rotation 45 in the patient
17 so that the fan beam 14 may irradiate a slice of the
patient 17 from a variety of gantry angles . The
radiation source 12 is controlled by a radiation control
module 48 which turns the radiation beam 14 on or off under
the control of a computer 51.

A compensator control 52 directed by a timer
generating desired position signals provides electrical
excitation to each electromagnet to control, separately,
the actuators 32 to move each of the leaves 30 in and out
of its corresponding sleeve 24 and ray 28 (see also Fig.
1). The compensator control 52 moves the leaves 30 of the
compensator 22 rapidly between their open and closed states
to either fully attenuate or provides no attenuation to
each ray 28. Gradations in the fluence of each ray, as
needed for each fluence profile, are obtained by adjusting
the relative duration during which each leaf 30 is in the
closed position compared to the relative duration during
which each leaf 30 is in the open position, for each gantry
angle.

The ratio between the closed and open states or the
"duty cycle" for each leaf 30 affects the total energy
passed by a given leaf 30 at each gantry angle and thus
controls the average fluence of each ray 28. The ability
to control the average fluence at each gantry angle permits
accurate control of the dose provided by the radiation beam
14 through the irradiated volume of the patient 17 by
therapy planning methods to be described below. The
compensator control 52 also connects with computer 51 to
allow program control of the compensator 22 to be
described.

A tomographic imaging system 11 employing an x-ray
source 46 and an opposed detector array 50 may be
advantageously mounted on the same gantry 44 as the
radiation source 12 to produce a tomographic or slice image
of the irradiated slice of the patient 17 prior to
radiation therapy for planning purposes. Alternatively,
such tomographic imaging may be performed on a separate
machine and the slices aligned according to fiducial points
on the patient 17.

A gantry control module 9 provides the signals
necessary to rotate the gantry 44 and hence to change the
position of the radiation source 12 and the angle  of the
fan beam 14 for the radiation therapy, as well as for the
computed tomography x-ray source 46 and detector array 50
also attached to gantry 44. Gantry control module 9
connects with computer 51 so that the gantry may be rotated
under computer control and also to provide the computer 51
with a signal indicating the gantry angle  to assist in
that control.

Control modules for the tomographic imaging system 11
include: x-ray control module 56 for turning on and off the
x-ray source 46, and data acquisition system 58 for
receiving data from the detector array 50 in order to
construct a tomographic image.

An image reconstructor 60 receives the data from the
data acquisition system 58 in order to assist in
"reconstructing" a tomographic treatment image from such
data according to methods well known in the art. The image
reconstructor 60 also communicates with computer 51 to
assist in high speed computations.

The tomographic
treatment image allows verification of the patient setup
just prior to radiation therapy treatment. An image
reconstructor 60 typically comprising a high speed array
processor or the like may use the actual fluence signals 57
and barrier signals 59 to produce a tomographic absorption
image to be used for verification and future therapy
planning purposes as described in more detail below.

A terminal 62 comprising a keyboard and display unit
63 allows an operator to input programs and data to the
computer 51 and to control the radiation therapy and
tomographic imaging equipment 10 and 11 and to display
tomographic images produced by the image reconstructor 60
on the display 63.

A mass storage system 64, being either a magnetic disk
unit or tape drive, allows the storage of data collected by
the tomographic imaging system 11 and the multi-segment ion
chambers 47, 53 for later use. Computer programs for
operating the radiation therapy system 10 will generally be
stored in mass storage unit 64 and loaded into the internal
memory of the computer 51 for rapid processing during use
of the system 10.

During operation of the radiation therapy unit 10, the
compensator control 52 receives from the computer 51 a
fluence profile for each gantry angle. The fluence profile
describes the intensity or fluence of each ray 28 of the
radiation beam 14 that is desired for that gantry angle 
at a given position of the patient support table (not
shown) as translated through the radiation beam 14. The
collection of fluence profiles over a range of rotation
gantry angles is termed a "treatment sinogram".

III. Therapy Planning Software

The generation of a treatment sinogram needed to
obtain the full benefits of the above described compensator
is performed by specially developed software running on the
computer 51 and reconstructor 60. Although the treatment
planning is performed in software, it will be recognized
that the planning may also be implemented in discrete
electronic circuitry dedicated to this operation and that
such dedicated circuitry may be employed to provide even
greater speed to this process.

Referring to Fig. 6(a), the generation of the desired
treatment sinogram to control compensator 22 begins with
the definition of a desired dose map 66. A typical desired
dose map 66 assigns a relatively high radiation dose,
within a dose constraint, to an area of tumorous tissue 68
and a second lower radiation dose to the area of healthy
tissue 70 outside of that region. The healthy tissue 70
may include an area 72 including a radiation sensitive
organ or the like to which an even lower radiation dose may
be assigned.

The desired dose map 66 is stored within the memory of
computer 51 as an array of elements each element holding
one digital value, and may be most easily entered by
displaying the tomographic view of the slice of patient 17
on the display 63 of the terminal 62 and manually tracing
around the tumorous area 68 using a track-ball or a similar
input device as is well understood in the art. Standard
area-filling computer programs may be used to transfer the
dose values assigned to each traced region to the
appropriate element in the array of memory representing the
desired dose map 66.

Each element of the dose map 66 thus defines the dose
desired at each of the plurality of volume elements 74
("voxels") within a slice of the patient 17. Referring to
Fig. 7, each voxel 74 of the patient 17 may be identified
by a vector r defined from a given reference point 76.
The dose at each voxel 74 is D(r).

A. Converting Dose to Terma1. Terma

Generally, the dose at any voxel r will depend on the
energy received at that voxel r from radiation scattered
from adjacent voxels (where adjacent voxels r include the
voxel r, i.e., the radiation received directly from the
radiation source 12). The dose D(r) for a given voxel r
is given by the following formula:
D(r)=∫T(r')A(r-r')d3r'
where T(r') is a value indicating the magnitude of the
primary total energy released from r' per unit mass of that
voxel r' and is called the "terma" (total energy released
per unit mass).

For a monoenergetic external radiation source, the
terma rate T ˙(r) is described by:
T(r')=µρ(r')E.(r')dt
where µ / ρ is an effective mass attenuation value at the
voxel r, E is the energy of the radiation photons in
Joules,  is the distribution of the fluence rate (flux
density). The integration of energy times fluence rate
over time is energy fluence Ψ(r) where:
Ψ(r')=E.(r')dt
hence
T(r')=µρ(r')Ψ(r')

Equation (4) basically relates how much energy from
the ray 28 interacts with the voxel r'.

2. Convolution Kernel

A(r-r') is a convolution kernel describing non-stochastic
energy transport or scattering in a uniform
medium. A(r-r') thus describes how the energy from each
voxel r' spreads to contribute to the dose at voxel r.

The kernel A(r-r') may be generated using a Monte
Carlo method as is generally understood in the art. As
mentioned, it is a three-dimensional function indicating
the fraction of energy absorbed at voxel r per unit of
energy released at voxel r'. The energy emitted from the
terma of each voxel r' finds its source in a directed ray
47 from external radiation source 12 and thus A(r-r') is
generally anisotropic as suggested in Fig. 8, spreading
outward away from the entry of ray 28. Energy conservation
requires that:
∫A(r')d3r' = 1.0

That is, if the energy transferred by the primary
interaction were all deposited on the interaction point,
the kernel would be approximated as a delta function.

Referring still to Fig. 8, the anisotropy of A(r-r')
is related to the gantry angle  and thus of the angle of
incidence of the ray 28 at r'. If the gantry angles  at
which the patient 17 is irradiated are predetermined, a
multi-direction convolution kernel B(r-r'), shown in Fig.
9, may be created from weighted superimposition of the
kernels A(r-r').

Referring to Fig. 9, assuming that the spreading of
radiation is approximately equal for all beam directions
and the rays 28 from each gantry angle  contribute equally
to the terma at voxel r', then the multidirectional
convolution kernel reduces to a "isotropic" form as
follows:
B(r-r') = 1ni=1nA(r-r')i
where n is the number of discrete gantry angles from
which rays 28 are projected.

For multiple rays 28 at different gantry angles, the
total dose at a given voxel r is the sum of doses from
each constituent beam, therefore:
D(r) = ∫ T(r')B(r-r')d3r'
where T(r') = nT(r')i, the latter term being the
contributed portion of the terma for the ith gantry angle.

This simplification assumes that the contribution to
the terma from each ray 28 is equivalent and takes
advantage of the distributive property of convolution.
Errors in this assumption are reduced by the filtration to
be discussed later.

Equation (7) substantially simplifies the calculation
of dose from terma but still requires a convolution for
each voxel r times the total number of voxels r' to
calculate the dose over the entire patient volume.
Preferably, therefore, the calculational efficiency of the
fast Fourier transform can be used and equation (7)
converted to the following:
D(r) = F-1{F{T(r')·F{B(r - r')}}
where F and F-1 symbolize Fourier and inverse Fourier
transforms respectively. This simplification of equation
(8) requires that the kernel B(r-r') be spatially invariant
and relies on the convolution theorem which states that
convolution of two spatially invariant quantities in a
space domain is equivalent to multiplication in the
frequency domain.

The assumption of the spatial invariance of B(r-r') is
correct only to a first order approximation. Typically,
the kernel B(r-r') for an external radiation source 12 is a
complex function of: (1) beam hardening of a polyenergetic
x-ray beam (i.e., the effect of the filtration of the
patient 17 in increasing the proportion of high frequency
or high energy radiation components), (2) the number of
rays 28 crossing each voxel, and (3) exponential
attenuation by the patient mass.

In the preferred embodiment, this first factor, beam
hardening, is neglected because it is an effect smaller
than the attenuation problem. Thus, the photon energy
spectrum in the patient 17 may be assumed to be the same as
that of the external radiation source 12. This
simplification is not required, however, and it will be
understood that beam hardening could be accurately
accounted for by representing a photon energy spectrum by a
finite number of separately convolved energy intervals.

The second factor, the difference in the number and
orientation of rays 28 that cross each voxel, caused by the
geometry of a finite number of gantry angles and the fan
orientation of the beam 14, affect spatial invariance.
Problems arising from the fan orientation of the beam (in
contrast to a parallel beam geometry) are largely solved if
there is a full rotation of the gantry 44. Errors
resulting from the fact that irradiation is performed at
only a finite number of gantry angles have been determined
to be acceptable.

The third factor affecting the assumption of spatial
invariance is the attenuation of the medium. This affects
the fractional contribution of the total terma from the
beams at each gantry angle. Accordingly, in those steps of
the planning procedure, as will be noted below, where
accurate calculation of dose is critical, the dose
distribution is calculated separately for each beam based
on the attenuation of overlying voxels, such attenuation
deduced from the parameters of the tomographic image. In
this case the simplification of equation (8) may not be
employed and repeated convolutions must be performed. In
certain steps in the planning process, however, as will be
noted, an estimate is sufficient and in these cases B(r-r')
is assumed to be spatially invariant and the dose
calculated according to equation (8).

Production of terma values from a desired dose map 75
is then simply the process of inverting equation (8) as
follows:
T(r') = F-1F{D(r)}F{B(r-r')}

This inversion requires that there be no significant
"zeros" (typically at high frequencies) in the denominator
term F{B(r-r')}or more simply that the kernel B(r-r') be
spatially compact (i.e., the Fourier transform of a
spatially compact kernel will have significant high
frequency content). It has been determined by the present
inventors that the kernels dictated for patients 59 are
sufficiently compact to allow this Fourier deconvolution.

Referring now to Fig. 10, this deconvolution to
produce a terma map 82, giving the terma for each voxel r,
from the desired dose map 66, is represented by process
block 80.

B. Converting Terma to Voxel Energy Fluence

Knowing the terma map 82, the energy fluence Ψ(r'),
which is a measure of the beam intensity, can be determined
at each corresponding voxel by equation (4) from a
knowledge of m/r as follows:
Ψ(r') = T(r')µρ(r')

The value of µ / ρ may be estimated and considered a
constant or actual m/r may be deduced from the tomographic
scan data collected by means of the tomographic imaging
system 60, (shown in Fig. 5). In this manner and as
illustrated by process block 84 of Fig. 10, a fluence map
86 giving the fluence at each point of the terma map may be
determined.

C. Converting Voxel Energy Fluence to Energy Fluence
Profile

The energy fluence Ψ(r') at each voxel r' is related
to the energy of the ray 28 exiting the compensator 22 by
the relation:
Ψ(r') = Ψ0(,)e-∫µρ(r)ρ(r)δ(p-r·r)dtSSD2(,)|t|2
where Ψ0(,) is the energy fluence for a given ray
28 as described by δ(ρ-r ∧·r) at the exit of the compensator
22 and serves to define the fluence profile of the
compensator and  and  are the gantry angle and the offset
angles of the ray 28 as previously described.

The exponential term represents the attenuation of the
ray 28 from the exit of the compensator 22 to the voxel r
caused by the mass of the patient 17 where by µ / ρ (r) is the
attenuation for each voxel r along the ray 28, ρ(r) is the
density of each voxel r, SSD(,) is the distance between
the exit of the compensator 22 and the surface of the
patient 17, r ∧ is a unit vector along r (where the origin
is now assumed to be the center of rotation of the gantry
44), and p is the perpendicular distance from the gantry's
center of rotation 45 and the ray 28. The vector is simply
a vector along the ray 28 to provide an integration
variable.

The fluence at each voxel r is related to the fluence
of the radiation beam 14 emitted from the compensator 22 by
equation (11). In the preferred embodiment, the density
and attenuation of each voxel r, µ / ρ(r) and ρ(r) are
assumed to be constant and the fan beam of radiation is
approximated by a parallel beam, hence
SSD2(,)|t|2 = 1.

Borrowing from the mathematics of tomographic image
reconstruction, the fluence map 86 may be "reverse" back
projected (i.e. projected) by projector 85 to determine a
fluence profile to be produced by the external-source
necessary to generate the desired fluence map and hence
dose.

This projection is simply the opposite of a typical
back projection used to form an image of a tomographic
slice of a patient 17 from a series of projections taken in
a tomographic imaging system. Because a projection is a
line integral across a distribution, the energy fluence
distribution for each voxel (equation (11)) is first
differentiated with respect to the rayline t:
dψ(r)dt = µρ(r)ρ(r)δ(ρ-r·r) = 2t

The line integral of dψ(r) / dt along t, corrected for
attenuation and inverse-square falloff, then represents the
projection operation and ψ0(,) the fluence profile over
the offset angles  of each gantry angle , is:
ψ0(,) = ∫µρ (r)ρ(r)δ(ρ-r·r) = 2tx(ψ(r)e+∫µρ(r)ρ(r)δ(ρ-r·r)dt|t|2SSD2(,)
↔(ρ-r·r)dt

The projection of equation (13) is represented by
projector 85 in Fig. 10.

The projection process, for the purpose of computing
fluence profiles for the compensator 22, differs in a
fundamental way from the simple inverse of tomographic back
projection. The difference is primarily in a concern for
the sharpness in the transition of the dose between the
irradiated tumorous tissue 68 and the healthy tissue 70.
Sharpness in this transition region reduces the irradiation
of healthy tissue 70 and is preferred over actual fidelity
of the dose to the desired dose map 66.

For this reason, the fluence map 86 from the fluence
calculator 84 is prefiltered as shown by process block 88
to produce a filtered fluence map ψ'(,) as follows:
ψ'(,) = F-1{F{ψ(,)|ω|}+
where ψ(,) is the fluence map 86 and |ω| is a ramp
filter in frequency space and the '+' subscript indicates
the positive component of the filtering result. This
prefilter 88 serves to increase the high frequency content
of the fluence map 86 and thus to aid in rapid transition
of dose at the tumor/non-tumor interface.

It is noted that this prefilter 88 is similar to the
filter used in tomographic imaging's "filtered" back
projection. That is, like tomographic imaging, the filter
de-emphasizes the low frequency components of the
projection in producing image data. In addition other
prefilters may be applied to correct for the use of the
radially symmetric kernel (equation (6)) in computing the
dose map from the terma map composed from the fluence
profile.

In practice the computation of a terma map, the
generation of a fluence map and the calculation of the
fluence profiles need not be performed as discrete steps
but may be accomplished by a direct projection of the dose
map with appropriate filtering. The filtering is
accomplished by a "fast inversion filter" which combines in
projection space the operation of deconvolution and ramp
filtration.

This may be symbolically specified by the following
equation

where

refers to a projection operation and I(t) is
the fast inversion filter. The ⊗ operators refers to a
convolution operation such as would normally be done in
Fourier space using a fast Fourier transformation.

Referring still to Fig. 10, the fluence profile
calculations of block 78, including the deconvolver 80, the
fluence calculator 84, the prefilter 88 which includes any
projection space filter (such as a ramp filter, a fast
inversion filter followed by truncation of zeros), and the
projector 85 thus produce fluence profiles which together
create an estimated treatment sinogram 87' from the desired
dose map 66. The fluence profile calculator 78 may use the
Fourier convolution of equation (9) in the estimate of the
fluence profiles at this stage, accepting minor
inaccuracies in that process, to be corrected at a later
stage, as will be described below.

D. Iteration

Referring now to Fig. 11, the fluence profile
calculator 78 converts the desired dose map 66 to an
estimated treatment sinogram 87', however the fluence
profiles of this estimated treatment sinogram 87' may not
be used to control the compensator 22 because, in general,
the estimated treatment sinogram 87' will include positive
and negative values of fluence. Only positive values of
fluence are physically realizable by the compensator 22; a
negative value of fluence would represent a ray 28 that
absorbed radiation along its path which is physically
unrealizable.

Accordingly, at process block 88, the fluence values
of the estimated treatment sinogram 87' are truncated to
positive fluence values 89. As a result of this
truncation, the estimated treatment sinogram 87' no longer
produces the desired dose map.

The amount of error resulting from the truncation
producing the positive fluence profiles 89 is determined by
back-projecting the positive fluence values 89 to an actual
dose map, 90 deviating from the desired dose map 66. This
back projection is accomplished by computing a fluence map
from the positive fluence values 89 per equation (11) and a
terma map per equation (4) and then convolving the terma
map with the kernel per equation (7) to establish the
actual dose map 90 per process block 92 of Fig. 11.

In this back projection, the assumption of spatial
invariance of the convolution kernel B(r-r') is not made so
as to produce a more accurate actual dose map 90.

The projection of a fluence profile onto a patient 17
to compute a dose map may be performed by a number of other
procedures known to those of ordinary skill in the art.

The actual dose map 90 is compared to the desired dose
map 66 to produce residual dose map 96 as indicated by
process block 94. In the preferred embodiment, the
comparison subtracts from the values of each voxel r of
the actual dose map 90, the greater of: a) the
corresponding value of desired dose map 66, or b) a
predetermined upper dose constraint. The predetermined
upper dose constraint is a threshold number that is deemed
an acceptable dose to tumorous tissue 76. Clearly, other
methods of quantifying the difference between the desired
dose map and the actual dose map will be apparent, from
this description, to those of ordinary skill in the art.

The result of this comparison process 94 is to produce
a residual dose map 96 shown in Fig 12(a). This residual
dose map 96 is then, again, operated on by the fluence
profile calculator 78 (in lieu of the desired dose map 66)
to produce an error fluence profile 98 (in lieu of the
estimated treatment sinogram 87).

A thus produced error fluence profile 98 is subtracted
by subtracter 100 from the estimated treatment sinogram 87'
to produce a new estimated treatment sinogram 90.

As shown in Fig. 11, this new estimated treatment
sinogram 87 is repeatedly operated on by process blocks 88,
92, 94 and 78 for a predetermined number of iterations, the
magnitude of the error fluence profile 98 values decreasing
with each iteration as shown in Figs. 12(b)-(c) until a
suitably low error fluence profile 98 is obtained.

The new estimated fluence profile 87 is then truncated
per process block 88 to produce a final sinogram 91 for use
in controlling the compensator 22, as previously described.

Referring again to Figs. 6(b), (c) and (d), dose maps
obtained by the present invention corresponding to a
desired dose map 66 of Fig. 6(a) are shown after: one
iteration (Fig. 6(b)); two iterations (Fig. 6(c)); and ten
iterations (Fig. 6(d)). The variations in dose in the
target volume shown in Fig. 6(d) are plus or minus 2% about
the predetermined upper limit of 1,000 cGy.

E. Converting Slice Data To Helical Data

In order to eliminate radiation hot spots and
radiation gaps along the length of a tumor, helical
scanning, in which the translation table is continuously
moved along a "z axis" through the gantry 44 as the gantry
44 rotates so that the radiation fan beam 14 sweeps a
helical pattern through the tumor site, is desirable.
Helical scanners also reduce irradiation time because the
start and stop motion of the translation table is
eliminated. Nevertheless, because of the constant
translation of the patient during helical scanning, the
treatment sinogram, as discussed above, must be modified.

Referring to Figs. 13(a)-(c), a simplified cylindrical
tumor 97 with its axis coincident with the axis of rotation
45 of the gantry 44 has been divided into separate slices
107, 108. An irradiation window 109 collimated by the
opaque mask 16 (shown in Fig. 1) subtends different
portions of the tumor 97 as the tumor 97 is translated
through the gantry 44 and the radiation source 12 rotates
about the tumor 97.

In a helical scan the irradiation window 109 subtends
multiple adjacent slices (i.e. it is not slice specific).
Referring to Fig. 13(a) at a 0∞ gantry position, the
irradiation window 109 may irradiate only slice 107. As
the tumor 97 is translated and the radiation source 12
rotated, the irradiation window 109 begins to subtend parts
of both slice 107 and adjacent slice 108. Referring to
Fig. 13(b), after 90∞ of rotation, the irradiation window
109 may subtend half of slice 107 and half of slice 108.
Referring to Fig. 13(c), after 180∞ of rotation, the
irradiation window 109 only subtends slice 108. Therefore,
helical or non-slice specific data must be developed from
the slice sinograms 91 in order to control the compensator
leaves 30 during helical irradiation.

Although different portions of a tumor may have
different cross sections or density distributions, if slice
data is generated that corresponds to many thin tumor
slices, the changes in the tumor between adjacent slices
will be small. Under these circumstances, fluence profiles
for adjacent tumor slices directed along identical gantry
angles  will be substantially similar. Therefore
interpolation between adjacent fluence profiles sharing an
identical gantry angles  may be made without appreciably
sacrificing irradiation accuracy.

To convert the final slice sinograms 91 to a
continuous fluence profile "ribbon" 93 for easy use in
helical scanning, a helical conversion module 95 (see Fig.
11) may use the following weighted averaging equation to
make adequate helical fluence profile approximations:
ψ(z1,z2,j) = ψ(z1,j)(1-j180)+ψ(z2,j)(j180)
where z1 is the first of two adjacent slices, z2 is a
second adjacent slice, j is the gantry angle, ψ(z1,z2,j) is
the fluence profile of the fan beam 14 directed so as to
subtend adjacent portions of tumor slices z1 and z2 from
gantry angle j, ψ(z1,j) is the fluence profile from the
final slice sinogram 91 corresponding to slice z1 from
gantry angle j and ψ(z2,j) is the fluence profile from the
slice sinogram 91 corresponding to slice z2 from gantry
angle j. j = 0° when the fan beam subtends only slice z1
and changes as the gantry rotates until the fan beam
subtends only adjacent slice z2 after 180° of rotation.
Therefore, at j = 0° (see Fig. 13(a)) only ψ(z1,j) will
influence ψ(z1,z2,j). At j = 90°, (see Fig. 13(b)) when
the translation table has moved the tumor 97 one half of a
slice thickness so that one half of the fan beam 14 is
directed at slice z2, one half of ψ(z2,j) and one half of
ψ(z1,j) will influence ψ(z1,z2,j).

After the fluence profile ribbon 93 is generated, it
is stored in the mass storage system 64 for later use
during a therapy session.

IV. Operation Of The Verification System

Prior to a therapy session, the fluence profile ribbon
93 for controlling the compensator 22 is loaded into the
compensator control 52. The ribbon 93 consists of a
plurality of fluence profiles to be directed at the tumor
from a sequence of gantry angles  as the translation table
moves through the gantry 44. Each fluence profile consists
of desired intensity data for each ray 28 of the fan beam
14. The compensator control 52, directed by the fluence
profile ribbon 93 data, drives the leaves 30 into and out
of the radiation beam 14 effecting various radiation
intensities as described above.

Referring to Figs. 3, 4 and 5, as the compensator
control 52 drives the leaves 30 between the open and closed
states, the trigger member 19 of each leaf 30 is driven in
and out of its associated light path 33. When the trigger
member 19 blocks the light path 33 (i.e. the leaf 30 is in
the closed state), the light detector 15 produces an actual
position signal indicating that the leaf 30 is in the
closed position. When the trigger member 19 is outside the
light path 33 (i.e. the leaf 30 is in the open state), the
light detector 15 produces an actual position signal
indicating that the leaf 30 is in the open state.

An error detector (not shown) realized in software
runs on the computer 51 and compares the actual position
signals to the desired position signals generated by the
compensator control 52 to identify mistakes in leaf 30
movement. If a leaf 30 fails to assume the position
indicated by a desired position signal, the compensator
control 52 assumes that the actuator 32 failed.

Upon failure of an actuator 32, the compensator
control 52 directs its signals so as to bypass the primary
actuator 32 and begin to direct the backup actuator 35.
Thus, a level of redundancy is added to the system wherein
failure of a leaf movement means will not require delay in
therapy protocol.

In the event that both actuators 32, 35 fail, the
spring 39 biases the leaf 30 into the closed state, the
leaf 30 thus occluding its associated ray 28. In this
manner the possibility of a leaf 30 becoming stuck in an
open state so that an uncontrolled ray 28 is directed at
the patient is reduced.

The system can be equipped with an alarm to indicate
when one or both of the actuators 32, 35 has failed.
However, even upon failure of both actuators 32, 35
associated with the same leaf 30, a therapy session can
continue once the leaf 30 is biased into its closed
position. A closed leaf 30 poses no danger of radiation
overexposure because the associated beam ray 28 is entirely
occluded. The deficiency in actual radiation absorbed by
the tumor because of leaf 30 failure can be compensated for
in later therapy sessions.

Referring now to Fig. 14, there is shown a simplified
compensator 22 and verification system 71 wherein the
compensator 22 has only eight attenuating leaves 30
dividing the fan beam 14 into eight adjacent rays 28.

As part of a treatment verification system 59, the
first multi-segment ion chamber 47 disposed between the
compensator 22 and the patient 17 has eight chamber
segments 49, each segment 49 directly within one of the
eight rays 28 of the fan beam 14. Each segment 49 produces
a measured ray fluence signal 55, or pre-patient fluence
signal indicative of the ray fluence encountered thereby.

The second multi-segment ion chamber 53 disposed
within the fan beam 14 opposing the radiation source 12 on
the opposite side of the patient 17 consists of a second
group of eight chamber segments 54, each segment 54
intercepting one ray 28

that passes through the patient
17. These chamber segments 54 produce post-patient fluence
signals 57 indicative of the fluence of rays 28' that have
traversed the thickness of the patient 17.

The measured ray fluences 55 are employed by the
verification system 71 for two distinct purposes. First,
referring to Fig. 15, a comparator module 101 receives from
the computer 51 the desired ray fluence to be generated by
every leaf 30 of the collimator 22. In addition, the
comparator module 101 (see Fig. 15) receives the measured
ray fluence 55 generated by the first ion chamber 49.
Comparing the desired ray fluence to the measured ray
fluence 55, the comparator module 101 generates a
difference value 102 for each leaf 30 at each gantry angle
.

A limit module 104 determines if the difference value
102 is outside of a reasonable limit and dangerous to the
patient. If so, the compensator control 52 turns off both
the primary 32 and backup 35 actuators for the duration of
the therapy session. With the actuators 32, 35 turned off,
the biasing spring 39 (see Fig. 4) forces the leaf into the
closed position to occlude the ray 28 and eliminate
possible radiation danger.

If the difference value 102 is minimal or the measured
ray fluence 55 is less than the desired ray fluence, the
difference can be corrected by adjusting the ray intensity
of second ray 28 and a different gantry angle.

Referring to Fig. 7, and as described above, the total
radiation absorbed by a voxel 74 of the tumor being the
summation of radiation directed along many rays 28 toward
the voxel 74, discrepancies between desired ray fluence and
measured ray fluence 55 can be corrected by adjusting the
fluence of the second ray 28 adjacent the first ray.
Decreasing fluence of a second ray 28 reduces the quantum
of total radiation delivered to a voxel 74. Likewise,
increasing fluence of the second ray 28 increases the
quantum of total radiation delivered to a voxel 74. In
this manner, relatively minor discrepancies between desired
and measured ray fluences can be eliminated resulting in
more accurate therapy sessions.

Second, the measured ray fluence signals 55 are used
in conjunction with the post-patient fluence signals 57 to
produce tomographic absorption images. By knowing the
fluence of each ray 28 entering the patient 17 and the
fluence of each ray 28' exiting the patient 17, a simple
subtraction calculation generates an absorption value
indicating how much radiation is absorbed by the tissue
within the patient 17 traversed by the ray 28. By
combining all of the rays 28 directed along a gantry angle
 while the translation table is in a single position, an
absorption profile for that angle  and table orientation
can be constructed.

Data collected during a helical therapy session in
which a fan beam 14 sweeps a helical pattern through a
tumor site is not slice specific. Because it is most
advantageous to view tomographic images as slices rather
than as a helix, the helical data is converted to slice
specific data.

The computer 51 employs the helical conversion module
95 for a second time to convert the helical data into slice
fluence data. Because each tumor slice shares absorption
profiles with two adjacent tumor slices (one before and one
after), the following weighted averaging equation may be
employed:
ψ(z2,j) = ψ(z1,z2,j)j+180180-1+ψ(z2,z3,j)360-j180
where z1 is a first tumor slice, z3 is a second tumor
slice, Z2 is a third tumor slice between z1 and z3, ψ(z2,j)
is the absorption profile for tumor slice z2 at gantry
angle j, ψ(z1,z2,j) is the absorption profile detected
between adjacent portions of slices z1 and z2 at gantry
angle j and ψ(z2,z3,j) is the absorption profile detected
between slices z2 and z3 at gantry angle j where j varies
between 0° and 360°.

After slice absorption profiles corresponding to the
various gantry angles  have been calculated for each tumor
slice, tomographic reconstruction techniques may be used to
produce a plurality of slice specific tomographic
absorption images to be viewed on the display unit 63.

Standard isodose curve data used in the art for
therapy planning purposes can be used by the computer 51 to
establish various levels of radiation absorption along the
depth of the tissue traversed by each ray 28.

By back-projecting a plurality of absorption profiles
(one profile for every gantry angle  at which the beam 14
was directed toward a slice) and simultaneously accounting
for the levels of radiation absorption within each ray 28,
a tomographic absorption image can be constructed in a
manner similar to that used in tomographic x-ray imaging.

A radiologist can use the tomographic absorption
images to determine radiation dose absorption within slices
of the tumor. These images can also be used to develop
more accurate diagnostic techniques and to study the
specific effects of irradiation on tumor size and
longevity.

Referring to Fig. 16, a second embodiment of a
radiation therapy unit 110 suitable for use with the
present invention includes a radiation source 112 producing
a generally conical radiation beam 114' emanating from a
focal spot 118 and directed towards a patient (not shown).
The conical beam 114' is collimated to form a generally
planar fan beam 114 centered about a fan beam plane 120.

As in the first embodiment, the compensator 122 is
centered in the fan beam 114 and about the fan beam plane
120, prior to the radiation being received by the patient.
However, the compensator 122 of the second embodiment
includes a rack 124 having an upper level 125 that extends
on both sides of the beam 114, well past edges 123 of its
thickness 115.

Referring to Fig. 17 the upper level 125 of the rack
124 has a set of sleeves 126 that slidably receive a
corresponding number of equispaced trapezoidal leaves 128.
Each leaf 128 is constructed of a dense, radio-opaque
material such as lead, tungsten, cerium, tantalum or a
related alloy.

The sleeves 126 are constructed of radio-translucent
materials and integrally attached to a mounting plate 136
which is fixed relative to the focal spot 118. The
mounting plate 136 is constructed of a sturdy radio-opaque
material and is positioned outside the fan beam 114 to
prevent interference with the fan beam 114.

The leaves 128, of the upper level together form an
arc of constant radius about the focal spot 118. When the
leaves 128 are in the closed position, a leaf gap 130
slightly less than the width of one leaf 128 exists between
each two adjacent leaves 128. The leaves 128 and leaf gaps
130 divide the fan beam 114 into a set of adjacent slab
like rays 138 at offset angles. The rays 138 that fall on
the leaves 128 are attenuated and those directed between
the leaves 128 are passed through the upper level 125
unattenuated.

Referring to Figs. 16 & 17, the rack 124 also includes
a lower level 132 that also extends past the edges of the
thickness 115 of the beam 114. Lower level 132 is disposed
between the upper level 125 and the patient. Like the
upper level 125, the lower level 132 has a second set of
sleeves 135 that receive a corresponding second number of
equispaced trapezoidal leaves 134. Leaves 134 are
constructed of the same material as leaves 128.

Each sleeve 135 of the second rack 132 is disposed
under a leaf gap 130 in the upper level 125 and is centered
within beam 114. Each leaf 134 in the lower level 132 is
sized to overlap any flanking leaves 128 above it so that
when viewed from the focal spot 118 it eliminates radiation
leakage. The overlap, however, of leaves 128 in the upper
level 125 and those in the lower level 132 is slight so
that very few rays 138 are attenuated by leaves in both
racks. The leaves 128, 134 of both levels are
substantially the same thickness, or at least thick enough
to entirely occlude their associated rays 138 when in the
closed position.

Referring now to Figs. 16 and 19(a)-(c), the sleeves
126, of both levels 125 are sized to permit each leaf 128
to slide to either side of the fan beam 114 completely out
of the path of the fan beam 114 yet to still be guided by
the sleeve 126. The sleeve 126, guides the leaf 128,
between a "first open state" (Fig. 19(a)), wherein the leaf
128 is positioned to one side of the beam 114 allowing
unobstructed passage of one ray 138 of the beam, and a
"closed state" (Fig. 19(b)) in which the leaf 128 is
positioned in the center of the sleeve 126 occluding one
ray 138 of the beam, and a "second open state" (Fig.
19(c)), in which the leaf 128 on the opposite side of the
beam 114 allowing unobstructed passage of the ray 138 of
the beam.

Referring now to Fig. 18, the leaves 128, 134 are
supported and guided within the sleeves 126, 135 by guide
rails 137 fitted into notches 142 cut along the edges of
the leaves 128, 134. The notches 142 allow the guide rails
137 to slidably retain the leaves 128, 134 within the
sleeves 126, 135 during motion between the two open and
single closed states.

Referring again to Fig. 16, each leaf 128 and 134 may
be moved rapidly between its open and closed states by
means of a corresponding pneumatic cylinder 139 connected
to the leaf 128 and 134 by a flexible link 140. The
pneumatic cylinders 139 have internal pistons (not shown)
that may be moved at high velocity between the ends of the
cylinders 139 by means of pressurized air coupled to the
cylinders 139 through supply hoses 141. The supply hoses
141 are fed by a compensator control (not shown in Figs. 16
or 17) to be described below. The pneumatic cylinders 139
are capable of applying high forces to the leaves 128, 134
to move them rapidly and independently between the closed
and both open states.

As best seen in Figs. 16 & 17, two solenoid stops 143
are associated with each leaf 128, 134. One pair of
solenoid stops 143 are disposed above each sleeve 126 on
the upper level 125, with one stop 143 outside each edge
123 of the beam thickness 115. Similar assemblies 143 are
disposed below each sleeve 135 on the lower level 132, one
outside each edge 123 of the beam thickness 115. The
solenoid stops 143 are used to ensure accurate positioning
of the leaves 128, 134 within the fan beam 114 in a manner
to be described in more detail below.

Referring to Fig. 20, each stop 143 has an electrical
solenoid coil 144 which, when energized, attracts a
concentrically located armature 145 against the force of a
spring 150. A flexible connector 147 connects the armature
145 to a stopping shaft 148 protruding from the stop 143 to
be retracted by the armature when the coil is energized.
Each solenoid stop 143 is securely fastened to its
respective level of the rack 124 by a mounting bracket 146.
A flange 149 on the armature 145 holds the spring 150
between itself and the underside of the coil 145 to bias
the stopping shaft 148 out from the stop 143 when the coil
144 is de-energized.

The flexible connector 147 is longitudinally rigid,
but laterally flexible to absorb the torque applied to the
stopping shaft 148 by impact with its associated leaf 128,
134 as will be described. The stopping shaft 148 is
constructed of a rigid but resilient material that will not
bend under the pressure of repeated collisions with a
collimator leaf 128, 134.

As seen in Fig. 19(b), when the solenoid coil 144 is
not energized, the spring 150 biases the armature 145 out
from the coil 144 which in turn forces the stopping shaft
148 through a shaft hole 153 in the rack 124, and into the
sleeve 126, into the path of leaf movement. When the
solenoid coil 144 is energized, the armature 45 retracts
into the coil 144 and out of the path of leaf movement.

Referring now to Figs. 19(a)-(c), during radiation
therapy, the stops 143 remain de-energized most of the time
with the stopping shafts 148 protruding into the sleeves
and thus locking their associated leaf 128 into one of its
three steady state positions, the closed position (19(b))
the first open position (19(a)) and the second open
position (19(c)).

When a leaf is to move from the first open position
(Fig. 19(a)) to the closed position (Fig. 19(b)), the
compensator control module 158 first energizes a first
solenoid stop 143a which moves a stopping shaft 148a up and
out of the sleeve 126 allowing the leaf to move toward the
closed position. High pressure air directed to the
pneumatic cylinder 139 then pushing the leaf 128 toward the
closed position (see Fig. 19(b)). The leaf 128 strikes the
second stopping shaft 148b where it is halted in the closed
position occluding the entire thickness 115 of the ray 138.
stopping shaft 148(b) ensures the leaf 128 is accurately
positioned. The first solenoid stop 143a is then de-energized
and the compressively loaded spring 150 expands
forcing the stopping shaft 148a down back into the sleeve
126. At this point, the stopping shafts 148a, 148b of both
solenoid stops 143a, 143b are within the sleeve 126 and
abutting the edges of the leaf 128. With both solenoid
stops 143a, 143b de-energized, the leaf 128 is secured in
the closed position against perturbing forces.

Similarly, to move the leaf 128 into the second open
position, the second stop 143b is energized and the
stopping shaft 148b is pulled out of the sleeve 126. The
leaf 128 is pushed to the second open position (see Fig.
19(c)). The second stop 143b is de-energized and the
stopping shaft 148b is forced into the sleeve 126 to lock
the leaf 128 out of the beam.

As will now be described, by alternating open states
from one side of the beam thickness 115 to the other, the
characteristic gradient attenuation along the thickness 115
of the ray 138 (i.e. the beam 114) can be eliminated.

Referring now to Fig. 21(a), when the leaf 128 is in
the closed state, entirely occluding the fan beam thickness
115 the initial cumulative intensity of radiation received
by the patient is zero. As the leaf 128 is moving to the
first open state (see Fig. 21(b)), the cumulative intensity
of the received radiation increases starting with the first
areas to be exposed. By the time the leaf 128 reaches the
first open state in Fig. 21(c), there is a gradient of
received radiation 121 across the beam thickness 115.

As the leaf 128 begins to return to the closed state
as in Fig. 21(d), the gradient attenuation 121 becomes
steeper until it is twice as steep as it was after the
first open cycle (see Fig. 21(e)).

In the next cycle, the leaf 128 is moved to the second
open state as in Fig. 21(f), the exact direction as the
previous cycle producing the opposite gradient 121. When
the gradients of the two cycles add together, the resulting
cumulative intensity 119 of the rays directed across the
beam thickness 115 becomes uniform as in Fig. 21(g).

Referring again to Figs. 19(a)-(c), it should be noted
that alternating movement of a leaf 128, results in a
uniform radiation exposure across the thickness of the beam
115 only if the velocity of the leaf 128 is constant as the
leaf 128 travels through the beam thickness 115. Without a
constant leaf velocity, a parabolic exposure profile across
the beam thickness 115 results as the exposure of
alternating open cycles is added.

While, practically, some acceleration of the leaf 128
as it travels through the beam widths will be acceptable,
ideally, most of the acceleration of the leaf 128 will take
place while the beam 114 is either fully occluded or
entirely unattenuated. Leaf acceleration can be limited to
these two conditions by placing the solenoid stop 143
outside the beam thickness 115 and configuring the leaf 128
to be wider than the beam thickness 115 so as to provide
for a first 133a and a second 133b acceleration gap on
opposite sides of the beam thickness 115.

Referring still to Figs. 19(a)-(c), when the leaf 128
moves from the first open state (Fig. 19(a)) to the closed
state (Fig. 19(b)), it is accelerated from a stationary
position to a constant velocity before the leading edge 129
of the leaf 128 passes the first acceleration gap 133a and
enters the beam thickness 115. As the leading edge 129
moves across the beam thickness 115 its velocity is
constant until it strikes the second stopping shaft 148b.
The leaf 128 is decelerated after the leading edge 129 is
within the second acceleration gap 133b and the following
edge 131 is within the first acceleration gap 133a.

In the closed position (Fig. 19(b)), the leaf 128
occludes the entire beam thickness 115 as well as both
acceleration gaps 133a, 133b. Moving from the closed state
(Fig. 19(b)), to the second open state (Fig. 19(c)), the
leaf 128 is accelerated from its closed stationary position
to a constant velocity before the following edge 131 exits
the first acceleration gap 133a. Restricting acceleration
and deceleration to the period when the leaf edges 129, 131
are within the acceleration gaps is also observed when
moving from the second open state to the closed state and
from the closed state to the first open state.

It has been found that the optimal acceleration gap
133 size is independent of acceleration potential and is
one fourth the beam thickness 115 (see Appendix).
Therefore, in order for the leaf 128 to occlude both the
beam thickness 115 and both acceleration gaps 133 when in
the closed state, the leaf 128 should ideally be
constructed 1.5 times as long as the beam thickness 115
(see Figs. 19(a)-(c)).

Using a leaf 128 that is wider than beam thickness 115
enables the stops 143 to provide a simple leaf deceleration
mechanism. When the leaf 128 is moved from its first open
state (Fig. 19(a)) to the closed state (Fig. 19(b)), upon
impact with the stopping shaft 148b, the leaf 128 may
temporarily bounce back off the stopping shaft 148b.
Piston pressure is continued to limit "leaf bounce" to
positions in which the impacting leaf edge 129 remaining
within its associated leaf gap 133b. Thus, leaf
deceleration will be limited to acceptable leaf 128
positions.

Preferably, this second compensator embodiment 110
employs the radiation therapy hardware discussed above in
connection with Fig. 5, and utilizes the therapy planning
software discussed above in connection with Figs. 10 and
11.

The above description has been that of a preferred
embodiment of the present invention. It will occur to
those who practice the art that many modifications may be
made without departing from the scope of the
invention. For example, the image reconstructor 60 could
produce a tomographic desired fluence image for comparison
with the tomographic absorption image to produce a
tomographic irradiation error image. The error image could
then be used to repeat the helical irradiation process to
correct for insufficient irradiation. Clearly the method
for planning radiation therapy is not limited to a
particular radiation source but may be used with any
radiation source which may be decomposed into separately
attenuated radiation rays. In addition, a computed
tomography system need not be incorporated with the
radiation therapy equipment but separate such equipment
could be used. The relationship between the terma values
and the fluence values may assume a constant density of the
patient to eliminate the need for a precise tomographic
scan of the irradiated area.

It should also be recognized that the pneumatic
cylinders 139 in the second embodiment could be constructed
to decelerate the leaves 128, 134 and limit the effect of
"leaf bounce" and leaf wear upon stopping shaft collisions
during leaf movement between states. Referring again to
Figs. 19(a)-(c), when leaf 128 is moved from the first open
state (Fig. 19(a)) to the closed state (Fig. 19(b)), once
the leading edge 129 is through the beam thickness 114, the
associated pneumatic cylinder 139 could produce a breaking
pulse that slows the leaf 128 before the leading edge 129
impacts the stopping shaft 148b. Similar deceleration
pulses could be utilized at the end of each leaf movement
to limit leaf and stopping shaft wear. In order to apprise
the public of the various embodiments that may fall within
the scope of the invention, the following claims are made:

Claims (9)

A radiation therapy machine having a radiation
source (12) for producing a radiation beam (14') directed
toward a patient (17) at a gantry angle, the beam
including a plurality of adjacent rays, and a compensator
having an attenuation means (32) disposed between the
radiation source and the patient (17) for independently
controlling the fluence of each ray of the beam (18'),
and a compensator control (52) for controlling the
attenuation means (22) according to first desired ray
fluences at a first gantry angle;
characterised in that:
the compensator further includes a pre-patient
monitor (47) disposed between the attenuating means (22)
and the patient (17) having a plurality of monitor
segments, adjacent monitor segments subtended by adjacent
rays of the beam (14'), each segment producing a fluence
signal proportional to the measured fluence of the ray
subtending the segment, comparison means (101) for
comparing the desired fluence of a ray to the measured
fluence to produce a difference value, and limit means
(104) for producing an error signal if the difference
value is outside of a predetermined error range.

A compensator according to claim 1 wherein the
limit means (104) also produces:

a high signal indicating that the measured fluence was
too high relative to the desired fluence yet within the
predetermined error range; and

a low signal indicating that the measured fluence was
too low relative to the desired fluence yet within the
predetermined error range.

A compensator according to claim 2 wherein the
compensator control (52) receives signals from limit means (104)
and communicates with the attenuation means (22) to:

reduce a second desired fluence at a second gantry
angle in response to receiving a high signal; and

increase a second desired fluence at a second gantry
angle in response to a low signal.

A compensator according to claim 1 wherein the
attenuating means (22) comprises:

a plurality of radiation attenuation leaves (30);

a supporting structure (24, 36) positioned generally between
the radiation source (12) and the patient (17) for guiding the leaves (30)
between a closed state within the radiation beam (14'), each leaf (30)
thus occluding one ray of the beam (14'), and an open state
outside of the radiation beam (14') to allow unobstructed passage
of the ray; and

motivation means (32) communicating with the control means
for independently moving each leaf (30) between the open and
closed states to effect an open to closed ratio producing
the desired fluence of each ray.

a closed state centered within the radiation beam (14'), the
leaf (30) thus occluding one ray of the beam (14');

a first open state with the leaf (30) displaced outside of
the radiation beam (14') on a first side of the one ray; and

a second open state with the leaf (30) displaced outside of
the radiation beam (14') on a second side of the one ray; and
wherein the motivation means (32) moves each leaf (30)
alternately between the first open and closed state and
between the closed and the second open state.

A compensator according to claim 5 wherein the
leaf (30) in the first open and second open states are spaced
away from the beam (14') by a predetermined acceleration
distance.

A compensator according to claim 6 wherein the
acceleration distance is equal to ¼ a beam thickness
measured along the path of movement of the leaf (30).

A compensator according to claim 7 wherein a
leaf width measured along the path of movement of the leaf,
is 1 ½ times as long as the beam thickness.

A compensator according to claim 4 wherein the
plurality of radiation attenuating leaves (30) includes a first
plurality of radiation attenuating leaves (128) and a second
plurality of radiation attenuating leaves (134) and wherein the
supporting structure (24, 36) includes a first supporting structure
for guiding the first plurality of leaves (128) between a closed
state, each leaf thus occluding every other ray of the
beam (14'), and an open state outside of the radiation beam and a
second support structure for guiding the second plurality
of leaves (134) between a closed state, each leaf thus occluding
those rays of the beam (14') not occluded by the first plurality
of leaves when the latter are in the closed state, and an
open state outside of the radiation beam wherein the first
plurality of leaves (128) in the closed state are positioned
closer to the radiation source than the second plurality of
leaves (134) in the closed state.